Magnetic resonance imaging (MRI) systems, for example, are frequently used in hospital and other medical facilities to scan and obtain images of biological tissues within a patient. By obtaining such images, a doctor or other medical professional can indirectly observe and carefully examine an anatomic region of interest (ROI) within a patient. Upon conducting such an examination, the doctor may then accurately diagnose a patient's malady and prescribe an appropriate treatment.
To successfully scan a region of interest within a patient, a modern MRI system typically includes a plurality of energizable superconductive magnet coils. When energized with electric current, these superconductive magnet coils are utilized to create small individual magnetic fields which cooperatively form a larger overall composite magnetic field (B0). By situating a patient within the magnetic field, one or more bodily regions of the patient may be scanned and imaged with supplemental help from RF transmit/receive coils and also gradient coils in the system. Although older MRI systems have traditionally incorporated resistive and permanent magnets, most modern MRI systems incorporate superconductive magnet coils with which larger magnetic fields and stronger field strengths are realizable for better imaging.
The superconductive magnet coils within a modern MRI system are generally constructed with superconductive wire. The superconductive wire itself is typically formed from a superconductive alloy material with which a characteristic “critical temperature” (TCRIT) is inherently associated. In general, when the temperature in an operating environment is above this critical temperature, the superconductive wire characteristically behaves in a somewhat resistive (i.e., “normal”) manner as it conducts electric current. When, on the other hand, the temperature in an operating environment is below this critical temperature, the superconductive wire characteristically behaves in a superconductive manner and is able to conduct electric current with almost no resistance. In general, the characteristic critical temperature inherently associated with a given type of superconductive wire is an extremely cold, cryogenic temperature. For example, a superconductive wire constructed with filamentary niobium-titanium (Nb—Ti) alloy windings has a characteristic critical temperature of about 10 K or −263° C. Thus, for such a superconductive wire to behave, operate, and conduct electric current in a preferred superconductive manner within a high-powered MRI system, the wire is situated in a super-cooled operating environment. Typically, in a modern MRI system, such a super-cooled operating environment is both established and maintained within a cryogen-filled vessel in a compartmentalized structure called a “cryostat.”
In addition to being constructed from superconductive wire, the superconductive magnet coils within a modern MRI system are also electrically connected in series with superconductive wire so as to generally form a closed-loop, magnet coil circuit. By largely situating the entire magnet coil circuit within a cryogen-filled vessel and maintaining the circuit at a super-cool temperature, the entire magnet coil circuit is thereby rendered superconductive. In such a superconductive state, electric current can be introduced into the closed-loop magnet coil circuit so that the electric current circulates through the loop and magnet coils in a substantially continuous fashion with very little resistance. In this way, the electric current persists in the loop with very little dissipation over time. For this reason, this superconductive loop with electric current persistently circulating therein is sometimes referred to as a “persistent loop.” By sustaining the persistent flow of electric current within the loop in this manner, small magnetic fields are respectively generated by the individual magnet coils as electric current passes through the coils. These small magnetic fields, as alluded to hereinabove, cooperatively form a larger composite magnetic field (B0) which can successfully be sustained for a period of time. In sustaining such a magnetic field, the scanning of one or more tissue regions within a patient is thereby facilitated so as to ultimately help produce images of the tissue regions.
To introduce and ramp up the flow of electric current in such a superconductive magnet coil circuit for operating a modern MRI system, an activatable heater is utilized to heat a small section of superconductive wire in the loop above its characteristic critical temperature (TCRIT). In this way, the small section of wire is rendered temporarily resistive. After such heating and successfully making a small section of the loop resistive, an electric power supply with resistive (i.e., non-superconductive) wire leads is then switchably connected to the loop in parallel with the small heated section. Upon being connected, the power supply is turned on to introduce and ramp up the flow of electric current through the magnet coils in the magnet coil circuit. Once the electric current has been ramped up to a desirable level, the heater is turned off. With the heater turned off, the small section of wire is quickly cooled down by the cryogens within the vessel that encloses the wire so that the small section of wire is again rendered superconductive. Once the small section of wire is superconductive, the electric current introduced into the loop is then able to flow virtually effortlessly through the section of wire. At this point, the output from the electric power supply may be slowly and gradually ramped down so that the electric current introduced into the loop is ultimately left circulating in a persistent manner through the magnet coils and the small section of wire of the overall superconductive magnet coil circuit. Once electric current is successfully introduced into the superconductive magnet coil circuit in this manner, the electric power supply may be switchably disconnected from the loop. In this way, the MRI system is made ready, at least in part, to scan a patient.
Sometimes, a modern MRI system must intentionally be shut down for disassembly, relocation, upgrading, replenishing burned-off cryogens, or performing general maintenance. To shut down the system, the electric current persistently circulating through the superconductive magnet coil circuit must generally be dissipated (i.e., “ramped down”) to zero. To ramp down the current, an activatable heater is utilized to heat a small section of superconductive wire in the persistent loop above its characteristic critical temperature (TCRIT). In this way, the small section of wire is rendered temporarily resistive. After heating and making a small section of the loop resistive in this manner, the electric energy within the loop may be transferred to an electric power supply that is switchably connected in parallel with the small heated section to thereby complete the shutdown. As an alternative or in addition thereto, the electric energy within the loop may also be dissipated via external diodes that are connected in parallel with the small heated section.
In emergency situations, it may be necessary to intentionally shut down an MRI system in a very short timeframe. To shut down the system, the electric current persistently circulating through the superconductive magnet coil circuit must generally be dissipated and ramped down to zero. To ramp down the current, an activatable heater is utilized to heat a section of superconductive wire in the persistent loop above its characteristic critical temperature (TCRIT) to thereby make the wire resistive. As the resistive section of the wire progressively “grows” and enlarges while being heated in this manner, the overall loop rapidly becomes extensively resistive as well. As a result, the electric energy within the loop is soon dissipated to zero. This process of heating and thereby making some portion of the persistent loop resistive in order to quickly ramp down the current is commonly referred to as “quenching.” During this process of quenching, the electric current must be carefully ramped down in a very controlled manner. The reason for such is because as the electric current is quenched, the electric and/or electromagnetic energy of the magnet coil circuit is thereby converted into significant amounts of thermal energy (i.e., heat) which is released. As this thermal energy is released and emanates from the loop circuit, surrounding liquid cryogens (for example, liquid helium) within the cryogen-filled vessel of the cryostat are commonly burned off in the form of a gas (for example, helium gas). In this gaseous form, the cryogens must either be recondensed for re-cycling and re-use or instead be altogether evacuated though a vent and later replenished. In light of such, therefore, if the electric current is ramped down and quenched too quickly or abruptly, large uncontrolled amounts of heat may suddenly be generated and emitted from the loop which could burn and permanently damage parts of the MRI system, including one or more sections of the loop itself and its magnet coils, or even other surrounding hardware. Any such permanent damage can be costly. Thus, for this reason, precision control and care must be exercised when intentionally quenching electric current in the superconductive magnet coil circuit of a modern MRI system.
Sometimes, electric current persistently circulating through the superconductive magnet coil circuit may suddenly be quenched by accident due to an unanticipated thermal disturbance which causes one or more sections of the persistent loop to become resistive. For example, in one possible scenario, mechanical stress and frictional movement of one or more superconductive magnet coils during operation may cause localized frictional heating which, if significant enough, may cause rapid and uncontrolled quenching of the electric current in the circuit. In another possible scenario, if liquid cryogens within the cryostat are not properly replenished to maintain the overall magnet coil circuit at a sufficiently cold temperature below its characteristic critical temperature (TCRIT), one or more localized resistive “hot spots” may develop in the persistent loop which may cause rapid and uncontrolled quenching of the electric current in the magnet coil circuit. In either such scenario and in many possible others, such rapid and uncontrolled quenching is likely to cause the release and emission of significant amounts of thermal energy (for example, up to 80 mega-joules (MJ)) in the form of heat. As a consequence of such thermal energy being “dumped” in this manner, parts of the MRI system may be burned and permanently damaged, including one or more sections of the loop itself and its magnet coils, and even surrounding hardware. Hence, as with intentional quenching, accidental quenching can be damaging and costly as well.
In the past, per convention, superconductive magnet coil circuits in MRI systems have largely been constructed as single closed-loop circuits. With, however, the push in recent years for MRI systems to accommodate higher electric currents and voltages as well as generate stronger magnetic fields to facilitate improved scanning, such conventional magnet coil circuits are now somewhat more prone to suffer the deleterious effects of excessively high voltage levels between adjacent layers of coil windings and/or localized heat damage, especially if electric current in a circuit is quenched in a somewhat uncontrolled manner within only one or a few particular sections of the circuit. For example, if one of the superconductive magnet coils in a single closed-loop superconductive magnet coil circuit suddenly begins to quench, a significant amount of thermal energy will be released as electric current is dissipated while passing through the coil when the coil is in such a normal (i.e., resistive) state. In addition, as the electric current is dissipated via the coil, the strong magnetic field current-generated in that same coil will collapse, thereby dumping energy in the form of heat onto that very same coil. As a result of such a highly localized release and dumping of energy, a burning “hot spot” may develop and cause permanent damage to the coil before all of the energy in the circuit has been released and dumped. Given such single-looped magnet coil circuits' susceptibility to being damaged during a quench, various combinations of quench-inducing resistive “heaters,” energy-dissipating “dump” resistors, and/or voltage-regulating diodes are now being strategically connected in parallel with (i.e., “shunted” across) the magnet coils situated in the magnet coil circuits of more recent and modern MRI systems. By connecting such various circuit elements within a magnet coil circuit in this manner, the magnet coils themselves are thereby situated and protected within multiple looped sections of the overall magnet coil circuit. Situated and protected as such, any quenching in one section of the circuit will, by design, trigger a progressive and cascading chain reaction of quenching in numerous other sections of the overall circuit. In this way, quenching of the electric current in the circuit is not effected via merely one or a few hot spots in the circuit. Instead, quenching of the electric current is effected via numerous sections in the circuit so as to more evenly distribute the releasing and dumping of thermal energy from the circuit. As a result, highly localized quenching via one or a few hot spots, which may permanently damage a magnet coil circuit, is largely prevented.
In general, to facilitate accurate scanning and imaging on an MRI system, the overall magnetic field produced by the system should be as homogeneous (i.e., uniform) as is possible. A conventional single-looped superconductive magnet coil circuit typically has multiple smaller-sized superconductive magnet coils that are spaced apart and electrically connected in series within the circuit instead of having only one or a few larger-sized magnet coils spaced apart and connected within the circuit. In this way, by strategically spacing the individual smaller magnet coils apart, a somewhat homogeneous magnetic field is generated whenever the magnet coils are energized with electric current. For improved field homogeneity, it has over the years become desirable to also establish one or more additional superconductive loop circuits of energized superconductive “secondary” coils proximate to the “primary” or “main” superconductive magnet coil circuit. In this way, by sizing the superconductive secondary coils and situating them proximate to the superconductive magnet coils of the main magnet coil circuit in a strategic fashion, the magnetic fields generated by the secondary coils interfere with and/or supplement the magnetic field generated by the main magnet coil circuit in a complementary fashion so as to help create an overall magnetic field (B0) which is more homogeneous and desirable for imaging in a particular sized volume of space. This method of “field correction” by adding one or more loop circuits of energized secondary coils to help smooth out any inhomogeneities in the magnetic field created by the magnet coils of the main magnet coil circuit is called “shimming.” For this reason, such secondary coils are commonly referred to as “shim coils,” and a secondary loop circuit within which they are electrically connected is commonly referred to as a “shim coil circuit” or simply a “shim circuit.” Although there are other different types of shim coils and shimming devices such as, for example, active resistive shim coils, gradient coils, and passive shim plates, superconductive shim coils are often highly desirable for their ability to shim magnetic fields for large spatial volumes.
In a conventional single-looped superconductive magnet coil circuit, when quenching of the electric current commonly flowing through the superconductive magnet coils of the circuit is initiated, the magnetic fields generated by the magnet coils then proceed to collapse. As these magnetic fields collapse, a tensive fluctuation in the level of current commonly flowing through the magnet coils is briefly induced. After such a brief tensive fluctuation, the level of current commonly flowing through the magnet coils then quickly settles down to zero, thereby concluding the quench event. As acknowledged in Lenz's law, such an induced fluctuation in the current level is due to the magnet coil circuit's attempt to conserve flux in response to the collapse of the magnetic fields in the magnet coils. In addition to the superconductive magnet coil circuit itself, if the superconductive shim coils of a single-looped superconductive shim coil circuit are situated sufficiently close to the superconductive magnet coils of the magnet coil circuit during the magnetic fields' collapse, a tensive fluctuation in the level of electric current flowing through the shim coils may be briefly induced as well. Briefly inducing a tensive current fluctuation in proximately situated shim coils in this manner is due to “inductive coupling” resulting from the electromagnetic phenomenon known as “mutual induction.” In general, the level or strength of mutual induction between any two proximately situated coils is largely dependent on (1) the relative distance between the two coils, (2) the respective physical dimensions of the two coils, and (3) the permeabilities of the two coils' respective cores. If such a tensive current fluctuation induced in the superconductive shim coils is large enough to cause the current in the superconductive shim coil circuit to reach a “critical level” at which quenching and the consequential uncontrolled release of large amounts of thermal energy is initiated, one or more highly localized hot spots may be created in the shim coil circuit. As a result, the shim coil circuit may inadvertently be permanently damaged.
To prevent a proximately situated superconductive shim coil circuit from being damaged during a quench event in a conventional single-looped superconductive magnet coil circuit, both the physicalities of the shim coils and the spacings of the shim coils relative to the magnet coils are typically optimized and adjusted to simply reduce inductive coupling between the magnet coils and the shim coils. Such optimization and adjustment is generally not overly difficult, for when the magnet coil circuit is conventionally configured as a single-looped circuit, the level of electric current flowing through the serially connected magnet coils in the circuit is the same and common to all magnet coils therein. Thus, when quenching occurs in a conventional single-looped superconductive magnet coil circuit, the electric current in the circuit initially fluctuates and thereafter ultimately settles down to zero in a uniform manner through all of the magnet coils serially connected in the circuit. Given such commonality and uniformity of current flow through the magnet coils of a conventional single-loop superconductive magnet coil circuit, design of a complementary superconductive shim coil circuit that will not be strongly coupled to the magnet coil circuit during quenching is rather straightforward, for the magnet coil circuit generally need only be singly addressed as a whole for shim circuit design purposes. Ultimately, therefore, by anticipating the circuit behavior and uniformly fluctuating electric current level of a given magnet coil circuit during a quench event via computer simulation, an appropriate shim coil circuit for functionally complementing the magnet coil circuit can be designed so as to minimize inductive coupling during any such quench event. In this way, any electric current induced in the shim coil circuit during a quench event in the magnet coil circuit is largely limited and prevented from reaching a level that might cause permanent damage to the shim coil circuit.
With, however, the above-described recent advent of a multi-section protected superconductive magnet coil circuit wherein magnet coils are situated and protected within multiple looped sections of the circuit, addressing a magnet coil circuit as a whole for shim circuit design purposes is generally no longer practical or appropriate. In particular, given the parallel circuitry of various resistive elements and diodes which uniquely define the multiple protected sections within such a magnet coil circuit, the levels of electric current respectively flowing through the individual magnet coils at a given point in time are not always the same during a quench event in the magnet coil circuit. That is, depending on the particular section(s) of the magnet coil circuit in which the quench event is initiated and the unique chain reaction design scheme hardwired therein for triggering both the shunting of current away from certain magnet coils and the inducing of quenching in certain other magnet coils to more evenly distribute the release of thermal energy from the circuit, the levels of current respectively flowing through the magnet coils connected within the various sections of the magnet coil circuit are likely to be different and non-uniform. Hence, by design, different current fluctuation levels, fluctuation times, and rates of ultimate current collapse within the various protected sections of the magnet coil circuit are likely to be realized during a given quench event. Thus, for any such given multi-section protected superconductive magnet coil circuit, it is generally no longer appropriate to simply address the magnet coil circuit as a whole for shim circuit design and damage protection purposes, for the behavior of such a magnet coil circuit is quite multi-faceted during a quench event. For this reason, a shim coil circuit intended to functionally complement such a magnet coil circuit should ideally be designed to address the multi-faceted and section-specific behaviors of the various protected sections within the magnet coil circuit. In this way, one or more sections of the magnet coil circuit do not unexpectedly induce an excessively high and potentially damaging level of current in the shim coil circuit.
In light of the above, there is a present need in the relevant art for a shim coil circuit or system that (1) successfully shims a magnetic field generated by a multi-section protected superconductive magnet coil circuit in an MRI system and also (2) successfully avoids damage caused by mutual induction during a quench event in a multi-section protected superconductive magnet coil circuit of an MRI system.
In addition to the above, there is also a present need in the relevant art for an external interference shield (EIS) coil circuit or system that (1) successfully shields a magnetic field generated by a multi-section protected superconductive magnet coil circuit in an MRI system from external disturbances and also (2) successfully avoids damage caused by mutual induction during a quench event in a multi-section protected superconductive magnet coil circuit of an MRI system.
In further addition to the above, there is also a present need in the relevant art for other somewhat analogous secondary coil circuits or systems that successfully avoid damage caused by mutual induction during quench events in the multi-section protected superconductive magnet coil circuits of, for example, nuclear magnetic resonance (NMR) spectroscopy systems, mass spectrometer systems, and the like.